Diffuse reflectance spectroscopy (DRS) is sensitive to the absorption and scattering properties of biological molecules in tissue and therefore can be used as a noninvasive in vivo tool to obtain quantitative information about the physiological and morphological properties (e.g., biomarkers) of human tissue. Thus, DRS can be utilized specifically as a diagnostic tool to detect various diseases that alter human tissue properties. Potential clinical applications of DRS include precancer detection and cancer diagnostics, intraoperative tumor margin assessment, and monitoring of tumor response to chemotherapy. Fiber optic probes are commonly used to deliver illumination light to, and collect diffusely reflected light from, a tissue specimen for DRS measurements. However, in order for DRS to be utilized in a clinic, frequent calibration is typically required to correct or compensate for a number of factors, such as lamp intensity fluctuations, wavelength-dependent instrument response, interdevice variations, and fiber bending losses that occur while a measurement is taken.
Calibration techniques presently used by biophotonics researchers typically rely on measurements using power meters, reflectance standards, and/or tissue phantoms (i.e., models that simulate human tissue and blood vessels). These calibration procedures are usually performed after the clinical measurements are completed. One particular DRS calibration method involves a two-step calibration procedure that utilizes the measured spectra of a spectrally flat diffuse reflectance standard (i.e., a reflective Spectralon puck) and a phantom of known optical properties in order to obtain the absolute reflectance spectra of a tissue sample. For example, a calibrated reference phantom spectrum is obtained by dividing the collected phantom spectrum with the collected puck spectrum. Similarly, a calibrated tissue spectrum is obtained by dividing a collected tissue spectrum with a second collected puck spectrum (i.e., a 2nd spectrum measurement of the same calibration puck). More specifically, calibration is performed by dividing the tissue spectra point by point by the spectra of the puck. Afterwards, a ratio of the calibrated tissue spectrum and the calibrated reference phantom spectrum is input into an inverse Monte Carlo model, which in turn extracts the optical properties of the tissue.
The aforementioned calibration of the tissue spectrum against a reference phantom is needed to put the experimental and Monte Carlo simulated data on the same scale. This is typically necessary no matter what type of calibration method is employed. However, the calibration of the tissue spectra and reference phantom spectra to the puck spectra is carried out to account for day-to-day system variations that occur between the time of the tissue measurement and the time of the reference phantom measurement.
There are a number of limitations associated with spectral data calibration methods currently utilized. Notably, these calibration methods fail to correct or compensate for real-time system fluctuations, such as variations in lamp intensity. For example, a given DRS illumination source typically requires at least 30 minutes of warm-up time to prevent significant light intensity fluctuations. However, the 30 minute warm-up period can pose considerable unwanted delays in a clinical setting, such as an operating room. Remarkably, light intensity of a light source can change as much as 25% during the warm-up period and even 3% afterwards. These variations in intensity are significant considering a 5% change in light intensity can introduce approximately 20% error in the extraction of optical properties from a tissue sample.
Another problem that arises in optical spectroscopy is the error caused by bending the optical fibers of the probe. Sharp bending frequently occurs in clinical applications, such as endoscopy, where the fiber optic probe is manually handled. For example, bending the detection arm (all 200 μm fibers) of the probe to a diameter of 3 cm (three turns) causes 6% light intensity attenuation, while bending the probe even further to a diameter of 2 cm can cause 11% attenuation in light intensity. As mentioned above, a 5% change in intensity can result in approximately 20% error in extracted optical properties from a tissue specimen.
Traditional calibration techniques typically rely on measurements from tissue phantoms and/or a diffuse reflectance standard that are usually performed after the clinical measurements are completed. Although these traditional calibration methods are successful in correcting instrument throughputs and remove day-to-day system drifts, none of the calibration methods are able to correct for real-time lamp fluctuations and fiber-bending loss while the specimen measurement is made. Similarly, all traditional calibration methods require at least 30 minutes for warming up the light source and a time-consuming calibration test procedure that is separate from the collection of the tissue sample spectra. As indicated previously, the reduction of unnecessary delays or procedures is extremely desirable in a clinical setting.
Thus, there remains a need for an improved system and method for performing optical spectroscopy using a self-calibrating fiber optical probe.